This disclosure relates generally to medical imaging, and more particularly to systems and methods for reconstructing an emission activity image.
Emission Tomography (ET), for example, Positron Emission Tomography (PET) imaging, Single Photon Emission Computed Tomography (SPECT) imaging, and the like, provides in-vivo functional information on metabolic and biochemical processes. In ET imaging, a solution including a radioactive tracer is injected into a subject (e.g., a human patient) to be scanned. Typically, the tracer moves towards and is taken up in one or more organs of the subject in which these biological and biochemical processes occur. Once a radioisotope of the tracer decays, it emits a positron, which travels a short distance before annihilating with an electron. The annihilation produces two high-energy photons i.e. gamma photons propagating in substantially opposite directions.
ET imaging uses a photon detector array arranged around a scanning area, usually in a ring-shaped pattern, in which the subject or at least a part of interest of the subject is arranged. When the detector array detects the two gamma photons within a short timing window, a so-called “coincidence” is recorded. A line connecting the two detectors that received the photons is called the Line Of Response (LOR). The reconstruction of an ET image is based on the premise that the decayed radioisotope is located somewhere on the LOR. Each coincidence may be recorded in a list by three entries, wherein two entries represent the two detectors and one entry represents the time of detection (also referred herein as emission projection data).
Based on the emission projection data acquired from an emission tomography scanner that uses the photon detector as described herein above, an emission activity image is reconstructed. The emission activity image represents the spatial and/or temporal distribution of radioactivity in the body and provides clinically useful information in oncology, cardiology or neurology applications.
One phenomenon that impacts the accuracy of reconstruction of emission activity image relates to attenuation of gamma photons traveling in the subject's body. For emission tomography to provide accurate quantitation, this attenuation needs to be appropriately taken into account during ET image reconstruction or data processing. Given a subject, an emission attenuation map, or simply an attenuation map, that represents a spatial map of linear attenuation coefficients for gamma photons (e.g., 511 keV photons for PET) is used for attenuation correction. The attenuation map obtained by some techniques and their limitations are described herein below.
A transmission scan using an external gamma photon source (e.g., 68Ge/68Ga source for 511 keV photons in PET) may be performed to extract information on attenuation. However, the transmission scan increases the scan time, and noise in the estimated attenuation map or attenuation correction factors is high due to low signal-to-noise ratio (SNR) transmission scan data. In addition, the emission tomography scanner needs to be capable of the transmission scan, and not every emission tomography scanner has the capability.
Alternatively, x-ray computed tomography (CT) scans, or simply CT scans, may be used for attenuation correction in emission tomography. X-ray CT scans are fast and may provide high SNR and high resolution CT images. In PET/CT (or SPECT/CT) scanners, no image registration is required between PET and CT (or between SPECT and CT) because PET and CT (or SPECT and CT) scanners are combined in a single gantry system. Conventionally, an x-ray CT image is reconstructed from CT projection data acquired from an x-ray CT scanner, or simply a CT scanner, and then the emission attenuation map for emission tomography is generated based on the reconstructed x-ray CT image. Since the x-ray photon energies (e.g., about 30-120 keV) are different from the gamma photon energy (e.g., 511 keV for PET), the x-ray CT image needs to be appropriately scaled to form the emission attenuation map. Conventionally, bi-linear or tri-linear scaling methods are used to scale the x-ray CT image in the Hounsfield unit in order to obtain the emission attenuation map. The scaling depends on the x-ray energy and the gamma photon energy, and needs to be calibrated for a CT contrast agent if any contrast agent is used.
Often, the trans axial field of view (FOV) of CT is smaller than that of emission tomography. In this case, the CT image is truncated or has truncation artifacts, and the emission attenuation map generated based on the CT image, accordingly, is truncated or has truncation artifacts. The truncation artifacts cause quantitation errors in the emission activity image. If a metal exists in the subject, metal artifacts may appear in the CT image and the metal artifacts are propagated into the emission attenuation map generated based on the CT image, resulting in quantitation errors in the emission activity image. As the CT dose is lowered or the number of CT views is reduced, artifacts may appear due to low dose or under-sampling in the CT image, and the artifacts may propagate into the attenuation map constructed based on the CT image.
Alternatively, the emission attenuation map and the emission activity image may be jointly reconstructed based on the emission projection data only. Such a method is called maximum likelihood reconstruction of attenuation and activity (MLAA), joint estimation or joint reconstruction. However, the problem of estimating both the emission attenuation map and the emission activity image from the emission projection data does not have a unique solution and the MLAA method usually suffers cross-talk artifacts. Although time-of-flight (TOF) information acquired from a TOF PET scanner alleviates the issues of non-uniqueness and cross-talk artifacts, it does not completely resolve them.